The present invention relates to a miniaturized, low power, programmable x-ray source for use in delivering substantially constant or intermittent levels of x-rays to a specified region. More specifically, the invention relates to apparatus and methods for delivering a uniform x-ray flux to an interior surface of a body cavity.
Most conventional medical x-ray sources are large, fixed position machines. Generally, the head of the x-ray tube is placed in one room and the control console in an adjoining area, with a protective wall, equipped with a viewing window, separating the two. The x-ray tube typically is approximately 20 to 35 centimeters (cm) long, and approximately 15 cm in diameter. A high voltage power supply is housed within a container located in a corner of the room containing the x-ray tube. Patients are brought to the machine for diagnostic, therapeutic, or palliative treatment.
Diagnostic x-ray machines are typically operated at voltages below 150 kilovolts (kV), and at currents from approximately 25 to 1200 milliamps (mA). By contrast, the currents in therapeutic units typically do not exceed 20 mA at voltages which may range above 150 kV. When an x-ray machine is operated at nominal voltages of 10 to 140 kV, the emitted x-rays provide limited penetration of tissue, and are thus useful in treating skin lesions. At higher voltages (approximately 250 kV), deep x-ray penetration is achieved, which is useful in the treatment of major body tumors. Super voltage machines, operable in the 4 to 8 megavolt (MV) region, are used to ablate or destroy all types of tumors, except superficial skin lesions.
A conventional x-ray tube includes an anode, grid, and cathode assembly. The cathode assembly generates an electron beam which is directed to a target, by an electric field established by the anode and grid. The target in turn emits x-ray radiation in response to the incident electron beam. The radiation absorbed by a patient generally is that which is transmitted from the target in the x-ray tube through a window in the tube, taking into account transmission losses. This window typically is a thin section of beryllium, or other suitable material. In a typical x-ray machine, the cathode assembly consists of a thoriated tungsten coil approximately 2 mm in diameter and 1 to 2 cm in length which, when resistively heated with a current of 4 amps (A) or higher, thermionically emits electrons. This coil is surrounded by a metal focusing cup which concentrates the beam of electrons to a small spot on an opposing anode which also functions as the target. In models having a grid, it is the grid which both controls the path of the electron beam and focuses the beam.
The transmission of an electron beam from cathode to anode is influenced by electron space charge forces which tend to become significant in conventional x-ray machines at currents exceeding 1 A. In such conventional machines, the beam is focused on the anode to a spot diameter ranging anywhere from 0.3 to 2.5 millimeters (mm). In many applications, most of the energy from the electron beam is converted into heat at the anode. To accommodate such heating, high power medical x-ray sources often utilize liquid cooling and a rapidly rotating anode, thereby establishing an increased effective target area, permitting a small focal spot while minimizing the effects of localized heating. To achieve good thermal conductivity and effective heat dissipation, the anode typically is fabricated from copper. In addition, the area of the anode onto which an electron beam is incident requires a material of high atomic number for efficient x-ray generation. To meet the requirements of thermal conductivity, effective heat dissipation, and efficient x-ray generation, a tungsten alloy typically is embedded in the copper.
In use, the total exposure from an x-ray source is directly proportional to the time integral of the electron beam. During relatively long exposures (e.g. lasting 1 to 3 seconds), the anode temperature may rise sufficiently to cause it to glow brightly, accompanied by localized surface melting and pitting which degrades the radiation output. However, thermal vaporization of the tube's coiled cathode filament is most frequently responsible for conventional tube failure.
While the efficiency of x-ray generation is independent of the electron beam current, it is highly dependent on the acceleration voltage. Below 60 kV, only a few tenths of one percent of the kinetic energy from an electron is converted to x-rays, whereas at 20 MV that conversion factor rises to 70 percent. An emitted x-ray spectrum is composed in part of discrete energies characteristic of transitions between bound electron energy levels of the target element. The spectrum also includes an x-ray energy continuum, known as bremsstrahlung, which is caused by acceleration of the beam electrons as they pass near target nuclei. The maximum energy of an x-ray cannot exceed the peak energy of an electron in the beam. Further, the peak of the bremsstrahlung emission curve occurs at approximately one-third the electron energy.
Increasing the electron current results in a directly proportional increase in x-ray emission at all energies. However, a change in beam voltage results in a total x-ray output variation approximately equal to the square of the voltage, with a corresponding shift in peak x-ray photon energy. The efficiency of bremsstrahlung radiation production increases with the atomic number of the target element. The peak output in the bremsstrahlung curve and the characteristic spectral lines shift to higher energies as the atomic number of the target increases. Although tungsten (Z=74) is the most common target material used in modern tubes, gold (Z=79) and molybdenum (Z=42) are used in some specialty tubes.
One disadvantage of most x-ray devices used for therapy is the high voltage, and consequent high energy radiation required when directed to soft tissue within or beneath bone. One example is in directing x-rays to areas of the human brain, which is surrounded by bone. High energy x-rays are required to penetrate the bone, but often damage the skin and brain tissue between the radiation entry site and the tumor. Another example in radiation therapy is in directing the x-rays to soft tissue located within the body cavity, couched among other soft tissue, or within an internal calciferous structure. Present high-voltage x-ray machines are limited in their ability to selectively provide desired x-ray radiation to such areas.
Another disadvantage of the conventional high voltage x-ray sources is the damage caused to skin external to the affected organ or tissue. Therefore, prior art high voltage x-ray sources often cause significant damage not only to the target region or tissue, but also to all surrounding tissue between the entry site, the target region, and the exit site, particularly when used for human tumor therapy. However, since present devices apply x-ray radiation to target regions internal to a patient from a source external to the target region, such incidental tissue damage is practically unavoidable.
Conventional radiation treatment of the soft tissue that lines body cavities, such as the bladder, vagina and cervix, urethra, uterus, colon and rectum, involves application of x-radiation from an extracorporeal source. Consequently, such techniques of radiation therapy have the disadvantage that they necessarily radiate areas of the patient between the radiation entry site, the target tissue, and the exit site, causing damages to such tissue.
Conventional methods of radiation treatment for body cavities also have the further disadvantage of failing to provide the ability to establish a uniform dose of radiation to the target tissue. In some cases, it is desirable that radiation treatment of the tissue lining a body cavity should provide the same dose of radiation to every segment of the tissue, i.e., a uniform, or other desired, dose. In other cases, specifically contoured non-uniform doses may be desired. The prior art x-ray sources cannot accomplish this for interior body cavities. As used herein, the term "uniform dose" refers to an isodose contour, i.e., a surface over which the flux density is substantially constant.
Some of these disadvantages can be overcome through the use of miniaturized low power x-ray sources, such as the one described in the above-referenced U.S. Pat. No. 5,153,900 issued to Nomikos et al. These sources can be inserted into, and activated from within, a patient's body. Thus, these sources can generate x-rays local to the target tissue. When such x-ray sources are used to treat the tissue lining a body cavity, the x-rays need not pass through the patient's skin, bone and other tissue prior to reaching the target tissue. However, even utilizing these sources there is no previously known method of providing a uniform, or other desired, dose of radiation to the target tissue, particularly where the geometry of the target region is not fixed, for example, as in the bladder which has a flexible inner wall without a well-defined shape.
By way of example, some miniature sources of the type disclosed in U.S. Pat. No. 5,153,900 generally act as point sources of x-ray radiation. Therefore, the strength of the radiation field decreases uniformly in air with approximately the square of the distance from the source (i.e., 1/R.sup.2). Since body cavities are not generally spherically symmetrical, a point source within a body cavity will not deliver a uniform dose of radiation to the tissue lining the cavity.
It is therefore an object of the invention to provide a method and apparatus for delivering a uniform dose of radiation to the tissue that lines a body cavity.
It is a further object of the invention to provide an apparatus, that includes a miniature low power x-ray source, for delivering a uniform or other desired dose of radiation to the tissue that lines a body cavity.
Other objects and advantages of the present invention will become apparent upon consideration of the appended drawings and description thereof.